In magnetic resonance imaging (MRI) applications, a patient is placed in a strong and homogeneous static magnetic field, causing the otherwise randomly oriented magnetic moments of the protons, in water molecules within the body, to precess around the direction of the applied field. The part of the body in the homogeneous region of the magnet is then irradiated with radio-frequency (RF) energy, causing some of the protons to change their spin orientation. This has the effect of nutating the net magnetization, which was directed with the static magnetic field prior to RF application, and thereby causing a component of the magnetization to be transverse to the applied static field. This precessing magnetization induces measurable signal in a receiver coil tuned to the frequency of precession (The Larmor frequency). This is the magnetic resonance (MR) signal. The useful RF components are those generated in a plane at 90 degrees to the direction of the static magnetic field.
The same coil structure that generates the RF field can be used to receive the MR signal or a separate receiver coil place close to the patient may be used. In either case the coils are tuned to the Larmor precessional frequency ω0 where ω0=γB0 and γ is the gyromagnetic ratio for a specific nuclide and B0 is the applied static magnetic field.
Conventionally, when imaging the thorax, a whole body radio frequency coil is used in both transmit and receive modes to enable full coverage of the anatomy. By distinction, when imaging the head, neck, knee or other extremity, local coils are often used as receivers in conjunction with whole-body transmitter coils. Placing the local coil close to the imaged region improves the signal-to-noise ratio and therefore the spatial resolution as well as limiting the field of view. In some procedures, local coils are used for both transmission and reception.
In some cases, a plurality of RF receiver coils forming an NMR phased array are used to enable MR signals from multiple regions in the body to be acquired at the same time (see for example U.S. Pat. No. 4,825,162). In this manner parallel imaging methods may be used to advantage in tailoring the region of interest and/or reducing scan times for comparable resolution to single receiver systems. Popular parallel imaging methods include “SMASH” (D K Sodikson and W F Manning, “Simultaneoaus acquisition of spatial harmonics (SMASH): fast imaging with radiofrequency arrays,” Magn. Reson. Med. 38:591-603, 1997) and “SENSE” (K P Pruessmann, M. Weigner, M B Scheidegger and P. Boesinger, “SENSE: sensitivity encoding for fast MRI,” Magn. Reson. Med. 42; 952-962, 1999).
In prior art phased array coils, the multiple receiver coils are arranged linearly in a plane, or can be wrapped circumferentially around a cylinder or similar shape. They are wrapped in a serial fashion, that is one coil after the other. The coils are placed either overlapping or adjacent to each other to eliminate their coupling. (See for example, J R Porter, S M Wright and A Reykowski, “16-element phased array head coil,” Magn. Reson. Med. 40: 272-279, 1998).
While these coils have been effective in producing complete images of an anatomical region, such as the brain, by combining signals from each of the array elements, it is a characteristic of all such prior art coil arrangements that the point of maximum sensitivity of each element is superficial to the anatomy under study. Often the area of diagnostic interest, in the head, for example, may be located away from the surface, deeper in the brain.
It is an object of the invention to provide an improved coil array in which each coil element has its maximum sensitivity close to the centre of the object under study.
It is a further object of the invention to provide a rotary switched phased array radio frequency structure.